U.S. patent number 5,052,394 [Application Number 07/328,082] was granted by the patent office on 1991-10-01 for method and apparatus for ultrasonic beam compensation.
This patent grant is currently assigned to Commonwealth Scientific and Industrial Research Organization. Invention is credited to David A. Carpenter, George Kossoff.
United States Patent |
5,052,394 |
Carpenter , et al. |
October 1, 1991 |
Method and apparatus for ultrasonic beam compensation
Abstract
In ultrasonic echoscopy, when an object to be examined has
overlying layers of a medium (for example, fat, muscle, skin or
bone) which have different ultrasonic transmission characteristics
from those of the object, the echogram of the object is usually
distorted. To reduce that distortion, the present invention uses a
transducer comprising an array of ultrasonic transducer elements
operated in a higher resolution mode than its normal imaging mode,
to obtain information about the geometry of the overlying layers.
Using this information and a knowledge of the transmission
characteristics of the overlying layers, amplitude and phase
corrections are calculated, to enable the transducer, when operated
in its normal imaging mode, to generate a required beam of
ultrasound. The corrections are then applied and an echogram of the
object, with reduced distortion, is obtained while operating the
transducer in its normal imaging mode.
Inventors: |
Carpenter; David A.
(Northbridge, AU), Kossoff; George (Killara,
AU) |
Assignee: |
Commonwealth Scientific and
Industrial Research Organization (Campbell, AU)
|
Family
ID: |
3772228 |
Appl.
No.: |
07/328,082 |
Filed: |
March 28, 1989 |
PCT
Filed: |
June 10, 1988 |
PCT No.: |
PCT/AU88/00181 |
371
Date: |
March 28, 1989 |
102(e)
Date: |
March 28, 1989 |
PCT
Pub. No.: |
WO88/09939 |
PCT
Pub. Date: |
December 15, 1988 |
Current U.S.
Class: |
600/442; 73/597;
73/599 |
Current CPC
Class: |
G01S
7/52049 (20130101); G01S 15/8918 (20130101); A61B
8/4494 (20130101) |
Current International
Class: |
G01S
7/52 (20060101); G01S 15/89 (20060101); G01S
15/00 (20060101); A61B 008/00 () |
Field of
Search: |
;128/660.01,660.06,660.10 ;73/597,599,602,625-626 |
References Cited
[Referenced By]
U.S. Patent Documents
Primary Examiner: Jaworski; Francis
Attorney, Agent or Firm: Ladas & Parry
Claims
We claim:
1. A method of generating a beam of ultrasonic energy for obtaining
an echogram of an object over which there is at least one layer of
a medium having different ultrasonic transmission characteristics
from those of the object, said method comprising these steps of
(a) positioning an ultrasonic transducer for examination of the
object, the ultrasonic transducer including an array of ultrasonic
transducer elements and being adapted to generate beams of
ultrasound at a fundamental frequency;
(b) operating the transducer in a higher resolution mode, in which
at least one of the transducer elements is operated at a frequency
which is higher than said fundamental frequency, to obtain
information about the geometry of the medium or media interposed
between the transducer and the object;
(c) calculating, using a knowledge of the ultrasonic transmission
characteristics of the medium or media and the geometrical
information obtained by step (b), the amplitude and phase
correction to be applied to the transducer elements of the
transducer to generate a required beam of ultrasonic energy at said
fundamental frequency within the object; and
(d) generating a beam of ultrasonic energy from the transducer by
operating the transducer at said fundamental frequency and applying
the corrections calculated by steps (c) to the phase amplitude of
the electrical pulses applied to activate the elements of the
array.
2. A method as defined in claim 1, in which the frequency used in
said higher resolution mode of said transducer which is higher than
said fundamental frequency is an overtone of said fundamental
frequency.
3. A method as defined in claim 2, in which said object is a
blood-carrying vessel and said method includes the additional step
of measuring the doppler frequency shift of the echoes from said
vessel when operating said transducer in its normal imaging mode in
step (d).
4. A method as defined in claim 1, in which said object is a
blood-carrying vessel and said method includes the additional step
of measuring the doppler frequency shift of the echoes from said
vessel when operating said transducer in its normal imaging mode in
step (d).
5. Apparatus for generating a beam of ultrasonic energy for
obtaining an echogram of an object over which there is at least one
layer of a medium having different ultrasonic transmission
characteristics from those of the object, said apparatus
comprising:
a) An array of ultrasonic transducer elements, each transducer
element being adapted to transmit ultrasound into the object when
activated and to receive reflected echoes of the ultrasound from
the object;
b) means to operate the array of transducer elements at a higher
resolution mode than its normal imaging mode, to obtain information
about the geometry of the or each overlying layer; and
c) means including a clock and delay lines to modify the normal
activation of the elements of the array when the array is used in
its normal imaging mode to apply a correction to the beam of
ultrasound generated by the array, calculated on the basis of the
information obtained about the geometry of the overlying layer or
layers, to reduce the distortion of the echogram image of the
object obtained using the array.
6. Apparatus as defined in claim 5, in which said means to operate
the array of transducer elements at a higher resolution mode
comprises means to operate at least one of the transducer elements
at a frequency which is an overtone frequency of the fundamental
frequency of operation of the array.
7. Apparatus as defined in claim 5, including means to measure the
doppler frequency shift of the echoes from the object when the
object is a blood-carrying vessel.
Description
TECHNICAL FIELD
This invention concerns ultrasonic echoscopy. More particularly, it
concerns a technique and apparatus for rectifying the degradation
of a beam of ultrasonic energy (ultrasound) as it passes through
layers of different media before reaching the object that is being
examined. Use of the invention will enable a significant
improvement in
a) cross-sectional ultrasonic images of the object being
examined;
b) the accuracy of measurement of parameters such as attenuation
and speed of propagation of ultrasound in human tissue using tissue
characterisation techniques;
c) the accuracy of measuring liquid flow in vessels using the
ultrasonic Doppler technique; and
d) the production of images and the taking of measurements within
the object being examined when through transmission techniques are
being used.
The invention is particularly, but not solely, directed to the use
of ultrasonic echoscopy techniques in medical diagnostic
examination.
BACKGROUND ART
Ultrasonic echoscopy provides information about an examined object
which may be displayed in the form of an ultrasonic echogram. Such
an echogram consists of a display of acoustic impedance
discontinuities or reflecting surfaces in the object. It is
obtained by directing a short pulse of ultrasonic energy, typically
having a frequency in the range of from 1 to 30 MHz, into the
object being examined. Any acoustic impedance discontinuities in
the object reflect and return some of the energy in the form of an
echo. This echo is received, converted into an electrical signal
and displayed as an echogram on a cathode ray oscilloscope, a film,
a chart or the like.
The echogram may constitute a one dimensional (A-mode) or a two
dimensional (B-mode) representation of the object being examined.
In both cases, the information is contained in the position and
magnitude of the echo displayed. In a one dimensional display, the
position along a base line is used to indicate the distance to the
reflecting surface whilst the magnitude of the echo is displayed,
for example, as a deflection of the base line or as an intensity
change. In a two dimensional display, the position along a base
line is used to indicate the distance to the reflecting surface (as
in a one dimensional display) and the direction of the base line is
used to represent the direction of propagation of the acoustic
energy. The two dimensional display is obtained by changing this
direction of propagation of the acoustic energy and by instituting
a similar, but not necessarily identical, movement of the base line
of the display. The magnitude of the echo is displayed as for a one
dimensional display, for example, as a deflection of the base line
or--more usually--as an intensity change.
The technique of ultrasonic echoscopy used in medical diagnosis to
obtain information about the anatomy of patients has been widely
reported. It has proved of particular value as a diagnostic aid
when examining the abdomen and uterus, eye, breast, lung, kidney,
liver and heart, these being regions of soft tissue with little
bone and air. In general, the technique is considered to compliment
other techniques to provide a more complete picture of a patient's
condition. However, particularly in pregnancies, ultrasonic
echoscopy may be useful in place of x-rays where the latter may not
give sufficient information or may be dangerous. In such medical
applications, a pulse of ultrasonic energy is transmitted into a
patient in a known direction and echoes are received from
reflecting surfaces within the body. The time delay between the
transmission of a pulse and the reception of an echo depends on the
distance from the transmitting ultrasonic transducer to the
reflecting surface. The distance information so obtained may be
displayed for interpretation and clinical use as a one dimensional
range reading or as a two dimensional cross section, as previously
described.
It is known (see, for example, the specification of U.S. Pat. No.
3,939,707 to Kossoff) to measure blood flow in the body along an
ultrasonic line of sight by measuring the frequency shift of
ultrasonic echoes and combining the frequency shift data with blood
vessel dimensional and directional information that has been
obtained from the B-mode ultrasonic echogram of the area. Using
this technique, an absolute measurement of the blood flow is
obtained.
It is assumed in the simple application of the pulse echo principle
that a pulse of ultrasonic energy propagates through the various
media of an object at a uniform velocity of propagation. In soft
tissues in a human body, this velocity of propagation is of the
order of 1570 meters per second (m/s). However, a scanning beam of
ultrasound will suffer distortion as it propagates through media
having different characteristics from those of the soft tissue
being examined. The distortion is due to refraction, which is a
consequence of the different velocity of propagation in each
medium, and attentuation (which is due to a number of effects,
including reflection at the interfaces between media and
scattering). The distortion is manifest as an unclear or inaccurate
echogram of the object being examined, and in deviation of the beam
from its initial line of sight, widening of the beam and an
increase in the level of the sidelobes of the transmitted
signal.
It has long been recognised that the removal of these distortions
will give an improvement in the resolution and clarity of the
resulting echograms. An improvement in the beam quality will also
improve the accuracy of tissue characterisation measurements and
will improve the signal to noise ratio for the ultrasonic Doppler
technique used to measure blood flow.
DISCLOSURE OF THE PRESENT INVENTION
It is an objective of the present invention to provide a method and
apparatus which can be used to reduce the distortions, due to
refraction and attenuation suffered by a beam of ultrasonic energy,
and thus provide improved echograms and improved measurements of
tissue characteristics and blood flow.
This objective is achieved by measuring the geometries of different
media overlying the region of interest in the body being examined,
and then using this information, and the known velocity of
propagation of ultrasound in such media, to construct a correctly
focused beam of ultrasound for use in the region of interest.
In the human body, for example, most organs are located behind
overlying tissues which are primarily skin, muscle and fat. Each of
these has a propagation velocity for ultrasound which is
significantly different from the normal velocity of propagation in
the internal organs, which is about 1570 m/s. In skin, the velocity
of propagation of ultrasound is about 1750 m/s, in muscle it is
about 1620 m/s, and in fat it is about 1440 m/s. The attenuation of
ultrasonic energy is also different in these media, with muscle
being more attenuating than fat. The practice of the present
invention involves the measurement of the geometry of the layers of
skin, muscle and fat (typically using high frequency ultrasound).
From this measurement of the thicknesses and shapes of these
layers, and with an a priori knowledge of ultrasound propagation
velocities and attenuations, it is possible to compute the required
ultrasound beam corrections in terms of time delays and signal
levels. These time delays and signal levels are then used to
construct a correctly focused beam of ultrasonic energy to
propagate deeper into the body, hence removing the distortions due
to the overlying layers.
The measurement of overlying layers can be achieved using an array
of ultrasonic transducer elements operated to achieve a high
resolution measurement of the object(s) being examined. There are a
number of known techniques which may be used to improve the depth
resolution of an ultrasound system, usually at the expense of
sensitivity or penetration depth in tissue (which is unimportant in
the present situation since only the geometry of the overlying or
superficial layers is being measured). Those techniques include
(this list is not exhaustive):
a) additional mechanical loading on the elements of the array via
changes in the backing layers and/or the matching layers;
b) changes to the electrical matching and loading of the
elements;
c) signal processing changes to give shorter pulses on transmission
and/or reception of signals; and
d) the operation of the array at a higher frequency (such as, at
the frequency of the third or fifth overtone).
The measurement of the geometry of the overlying layers can also be
performed using a single ultrasonic transducer element,
independently or in such an array, or using any number of elements
in such an array less than the full array.
After applying the necessary corrections, a beam of ultrasound
generated in the normal imaging mode by the array is used to obtain
an echogram of the deeper internal structure(s) being
investigated.
Thus, according to the present invention, there is provided a
method of generating a beam of ultrasonic energy for obtaining an
echogram of an object over which there is at least one layer of a
medium having different ultrasonic transmission characteristics
from those of the object, said method comprising the steps of
a) positioning an ultrasonic transducer for examination of the
object, the ultrasonic transducer including an array of ultrasonic
transducer elements and being adapted to generate beams of
ultrasound at a fundamental frequency (usually at 3.5 MHz);
b) operating the transducer in a higher resolution mode, in which
at least one of the transducer elements is operated at a frequency
which is higher than said fundamental frequency, to obtain
information about the geometry of the medium or media interposed
between the transducer and the object;
c) calculating, using a knowledge of the ultrasonic transmission
characteristics of the medium or media and the geometrical
information obtained by step (b), the amplitude and phase
corrections to be applied to the transducer elements of the
transducer to generate a required beam of ultrasonic energy within
the object; and
d) generating a beam of ultrasonic energy from the transducer by
operating the transducer at said fundamental frequency and applying
the corrections calculated by step (c) to the phase and amplitude
of the electrical pulses applied to activate the elements of the
array.
As noted already, one method for obtaining higher resolution is to
operate the transducer at a harmonic of the fundamental
frequency.
The present invention also encompasses apparatus for generating a
beam of ultrasonic energy for obtaining an echogram of an object
over which there is at least one layer of a medium having different
ultrasonic transmission characteristics from those of the object,
said apparatus comprising:
a) an array of ultrasonic transducer elements, each transducer
element being adapted to transmit ultrasound into the object when
activated and to receive reflected echoes of the ultrasound from
the object;
b) means to operate the array of transducer elements at a higher
resolution mode than its normal imaging mode, to obtain information
about the geometry of the or each overlying layer; and
c) means including a clock and delay lines to modify the normal
activation of the elements of the array when the array is used in
its normal imaging mode to apply a correction to the beam of
ultrasound generated by the array, calculated on the basis of the
information obtained about the geometry of the overlying layer or
layers, to reduce the distortion of the echogram image of the
object obtained using the array.
An embodiment of the present invention will now be described, by
way of example only, with reference to the accompanying
drawings.
BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1 illustrates, schematically, the generation of a focused beam
of ultrasonic energy by exciting an array of transducer elements,
as is known in this art.
FIG. 2 shows the operation of a long linear array transducer which
provides a linear scan of an area of interest, also as known in
this art.
FIG. 3 is a schematic sectional view of part of the human body,
showing a typical distribution of the layers of media, with an
ultrasonic transducer applied to the body and ultrasound beam
dispersion occurring at the interface between an overlying region
of muscle and an internal organ.
FIG. 4 is a similar view to FIG. 3, but with the ultrasound beam
corrections applied.
FIG. 5 is a block diagram of a modified linear array scanner that
may be used to practice the present invention.
DETAILED DESCRIPTION OF THE ILLUSTRATED EMBODIMENT
It is common practice in modern medical ultrasonic imaging (also in
radar and sonar imaging) to generate the imaging beam using an
array of transducer elements. In the arrangement shown in FIG. 1,
each of five ultrasonic transducer elements 11, 12, 13, 14 and 15
contributes to part of the overall aperture 10 used to transmit a
beam 17 of ultrasonic energy. Similarly, these transducer elements
make up the total aperture used to receive the beam reflected from
the target.
The linear array scanner illustrated in FIG. 2 consists of an array
20 of transducer elements 21 which sends out a focused beam of
ultrasound at right angles to the transducer face. The beam is
focused by applying time delays to the energising pulses as they
are sent to the array elements to produce the transmitted beam.
Such energising pulses are referenced 16 in FIG. 1. Similarly, time
delays are applied to the signals received on the transducer
elements 21 to produce a focused beam on reception. The beam is
scanned over the area of interest by switching to another group of
elements for each new line of sight. For example in the array of
FIG. 2, beam 22 is formed by the first five elements in the array,
beam 23 is formed by the second to sixth elements, beam 24 is
created by the third to seventh elements, and so on along the
transducer array.
In the present invention, the linear array forming the transducer
may need to be modified, as noted above, to operate in a higher
resolution mode, to measure the overlying layers. In the following
description of the illustrative embodiment, it will be assumed that
the array is to be operated at a higher frequency to increase the
measurement resolution. Thus, if the array of elements 11, 12, 13,
14 and 15 of the array 20, is a 3.5 MHz array, the transducer
elements may be excited at 10.5 MHz (the third overtone of 3.5
MHz). The beam of ultrasound produced at this higher frequency is
used to measure the distance to the interface between the overlying
tissues and an organ of soft tissue that is being examined. In the
example given in FIG. 3, this high frequency operation gives high
resolution images of the interface 31 between the muscle tissue
layer 32 and the liver 33.
The frequency of operation for imaging should be as high as
possible for good resolution. However, the attenuation of
ultrasound by tissue is proportional to the frequency of the
ultrasound, and a maximum frequency exists for each required
penetration depth. Since the interfaces are at a close range, a
considerably higher frequency can be used, thus affording a better
range resolution for the measurement of layer thickness. As noted
above, the high frequency mode of operation may use the normal set
of ultrasonic transducer elements used for imaging or a subset of
the elements, with different degrees of focusing, in order to
obtain the optimum resolution in measuring the geometry of the
overlying layers.
Having obtained a high resolution measurement of the distribution
of the layers overlying the organ of interest using the high
frequency higher resolution mode (or any of the other known higher
resolution procedures), the required velocities and amplitude
variations to correct the ultrasound beam characteristics when
operating in the normal imaging mode can be assigned.
It is necessary to translate the image corrections into time
differences to be imparted to various parts of the beam, and to
compute the difference in time delay which must be imparted to the
activating signal of each transducer element in the array of
transducer elements which generates the transmitted ultrasound beam
and receives the echoes from acoustic discontinuities, to
compensate for the different propagation velocities in the tissue
layers along a particular line of sight. Similarly it is necessary
to compensate for the different attenuations experienced by
different parts of the beam.
In the situation depicted in FIG. 4, the contributions to the beam
of ultrasonic energy which are provided by the elements 41 and 42
propagate through less muscle 32 than the contributions from the
elements 43, 44 and 45. Hence the contribution to the ultrasound
beam from the transducer elements 41 and 42 is attenuated less and
refracted more due to the higher propagation velocity in the muscle
32. This can be corrected by applying additional gain to the signal
transmitted (and received) by the transducer elements 43, 44 and 45
and extra time delays to the signals used to activate the elements
41 and 42, so that by the time the beam enters the liver 33, the
element contributions will be of the correct amplitude and in the
correct phases to produce a focused wavefront 40 and hence a
focused beam 46. With these new amplitude and time delays set, the
linear array can be operated along this line of sight at its
fundamental frequency (typically 3.5 MHz when imaging a human
abdomen).
If there are two or more interfaces (for example, a fat and muscle
interface followed by a muscle and organ interface), this process
can be repeated a number of times to compensate for distortions due
to the multiple media layers.
The correction process has to be repeated for each line of sight
along the linear array scan to obtain a good image of the area of
interest in the deeper structures. The correction process also has
to be repeated for each line of sight along different locations of
the array in the plane perpendicular to the paper containing FIGS.
3 and 4, if multiple echograms are required, for the thickness of
the overlying tissue varies in two dimensions.
As the measurement and compensation process described above can be
carried out quickly by the modified scanner, the imaging can
proceed at close to real time scanning speeds. The frame rate of a
real-time imaging system is dependent on the time required for the
desired number of ultrasonic pulses to propagate twice through the
maximum depth of tissue (that is, once to travel into the tissue
and once to return to the transducer). The high resolution
measurement step requires additional time but since the maximum
depth of penetration for the overlying layers is small, the
additional time is small and the overall increase in scanning time
is low.
FIG. 5 shows a partial block diagram of the apparatus needed to
perform the invention. The diagram shows the circuitry needed for
one element in the aperture. In the case of the transducer shown in
FIGS. 1 and 2, where five transducer elements are used for each
beam of ultrasonic energy, most of this circuitry would have to be
repeated five times. The exceptions are the clock 57, the
transmitter power supply 58, the adder 61, the time comparator 67,
the correction table 58 and the echo level measurement and
comparator circuitry 69, which are common to all transducer
elements in the aperture. The following components are basic to the
operation of a linear array ultrasonic scanner and are well known
in the ultrasonic echoscopy art:
a) the transmitter circuit 55 used to excite the array element at
its fundamental frequency (usually 3.5 MHz);
b) the transmit delay circuit 56 used to time the excitation of the
elements to produce a focused transmit beam;
c) the clock circuit 57 which produces the timing of transmit
pulses for the scanner;
d) the transmit power supply 58 that is used to set the output
level of the scanner;
e) the preamplifier 59 that operates to increase the echo level
received on the transducer element;
f) the receive delay line 60 used to set the timing of the returned
echoes to produce a focused receive beam;
g) the adder 61 which combines the signals from all the receive
elements; and
h) the switch 62 which connects the elements of the array needed to
form a particular beam to the appropriate electronics.
The switch 62 may connect different elements or combinations of
elements into the system when it is being used in the higher
resolution mode to measure the geometry of the overlying layers
from when the system is being used for normal imaging of an
object.
The following are the additional circuits that are needed to carry
out the method of the present invention:
i) the higher resolution transmitter 63 (which, in the
implementation described above, operates at a frequency of 10.5
MHz);
j) the switch 64 between the fundamental frequency transmitter 55
and the higher resolution transmitter 63;
k) an echo detector circuit 65 to detect the echoes from the
overlying tissue layers in the higher resolution measurement mode
of operation;
l) a time measurement circuit 66 which measures the time between
the transmit pulse and the echoes from the overlying layers;
m) a time comparator circuit 67 that is used to compare the times
measured in the time measurement circuit 66 between each of the
elements used in the aperture to form the beam;
n) a correction table circuit 68 which uses the times measured in
the time comparator circuit 67 to set the correct values of the
transmit delays provided by the delay line 56 and the receive
delays provided by the receive delay line 60, to give the correctly
compensated focused beam within the body (this can be implemented
by computing the corrected time delays or by using a look-up table
of corrected times);
o) an echo level measurement and comparator circuit 69 to measure
the echo level received on each element from the interface of the
overlying tissue layer and, by comparing levels between elements,
to set the correct transmit voltage levels for each element, thus
ensuring that each transducer element in the aperture will
contribute the correct level to form the compensated focused beam
within the body; and
p) a transmit voltage regulator circuit 70 to allow the levels set
by the echo level measurement circuit 69 to adjust the voltage
applied to the fundamental frequency transmitter 55.
With this identification of the circuit components, the operation
of the arrangement illustrated in FIG. 5 will be self-evident to
persons of skill in this art.
Those skilled in this art will also appreciate that although a
specific form of the present invention has been illustrated and
described above, modifications thereto may be made without
departing from the present inventive concept. For example, as
indicated earlier, different types of transducer arrays may be
used. In particular, each linear array of transducer elements that
has been included in the illustrated embodiment of the apparatus of
this invention may be replaced with an arcuate array of transducer
elements, which performs a mechanical pre-focusing of the
ultrasound beam. In such an arrangement, the position of the focus
of the beam is adjusted by varying the phase of the actuating
excitation pulses supplied to the individual transducer elements in
the array. Also, although a transducer having an array of five
transducer elements has been illustrated in FIGS. 1, 3 and 4, any
suitable number of transducer elements may be included in an array.
And, as indicated earlier in this specification, any of the
alternative known techniques for achieving higher resolution of the
images of the overlying layers may be used instead of the higher
frequency operating mode of the ultrasonic transducer.
* * * * *